Sound waves that have a frequency greater than approximately 20 kHz are referred to in the art as “ultrasound.” In the medical field, ultrasound waves are useful for both diagnostic and therapeutic applications. Medical diagnostic ultrasound systems are useful for generating images of anatomical structures within a target area of a patient's body. The images are obtained by scanning a target area with waves of ultrasound energy. In therapeutic ultrasound applications, high intensity ultrasound energy is transmitted into a target area to induce changes in state of the target. High-intensity focused ultrasound (“HIFU”) pulses induce changes in tissue state through thermal effects (e.g., induced hyperthermia) and mechanical effects (e.g., induced cavitation).
The use of high intensity focused ultrasound to destroy tissue or to alter the characteristics of tissue at a target location, volume, region or area within a larger mass, body or area of anatomical tissue presents many advantages, including minimization of trauma and pain for the patient, reductions in surgical incisions, stitches and exposure of internal tissue, avoidance of damage to tissue other than that which is to be treated, altered or removed, lack of a harmful cumulative effect from the ultrasound energy on the surrounding non-target tissue, reduction in treatment costs, elimination of the need in many cases for general anesthesia, reduction of the risk of infection and other complications, avoidance of blood loss, and the ability for high intensity focused ultrasound procedures to be performed in non-hospital sites and/or on an out-patient basis.
In high-intensity focused ultrasound hyperthermia treatments, intensity of ultrasonic waves generated by a highly focused transducer increases from the source to the region of focus where it can reach a very high temperature. The absorption of the ultrasonic energy at the focus induces a sudden temperature rise of tissue, which causes ablation of a target volume of cells in the focal region. Thus, as an example, HIFU hyperthermia treatments can cause necrotization of an internal lesion without damage to the intermediate tissues. The focal region dimensions are referred to as the depth of field, and the distance from the transducer to the center point of the focal region is referred to as the depth of focus. Ultrasound is a promising non-invasive surgical technique because the ultrasonic waves provide a non-effective penetration of intervening tissues, yet with sufficiently low attenuation to deliver energy to a small focal target volume. Currently there is no other known modality that offers noninvasive, deep, localized focusing of non-ionizing radiation for therapeutic purposes. Thus, ultrasonic treatment has a great advantage over microwave and radioactive therapeutic treatment techniques.
The beam emitted by a single ultrasound focused transducer element is generally effective within a fixed region, called the “focal zone.” This focal zone frequently is smaller than the size of the target tissue. Treatment of extensive targets is consequently a problem. A solution to this shortcoming is to utilize a transducer comprising a plurality of individual transducer elements arranged closely together to form an array. These arrays are focused at the desired treatment site through a combination geometric and electronic focusing. Geometric focusing is determined by the physical geometry of the array, while electronic focusing involves the use of phase delays and wave interference to achieve constructive interference at the target tissue. Electronic focusing allows movement of the treatment location without the need for mechanical movement of the array.
When a plurality of transducer elements are arranged in an array and are energized to propagate a steerable acoustic beam, the beam includes a main lobe, a plurality of small side lobes and one or more grating lobes. Grating lobes originate from acoustic waves that combine along various axes that differ from the axis of the main lobe. Grating lobes are present in the acoustic beams propagated from nearly all linear arrays and normally contain substantial energy. Side lobes are secondary points of focus outside the treatment region as a result of diffraction of the ultrasound waves passing through a structure, such as tissue or bone. The appearance of grating lobes and side lobes decreases the power radiated to the focal point, reducing the effectiveness of treatment.
It is known in the art to minimize grating lobes and side lobes by transmit apodization, wherein the effective drive amplitude of at least a portion of the elements of an array is varied to affect the shape of the beam. Varying the transmit amplitude at individual elements of the array typically requires the use of variable or switched-rail power supplies or inefficient linear amplifiers at each drive channel. Such drive mechanisms are slow, prohibitively bulky, and costly, particularly considering the large number of array elements present in a typical transducer and the number of element apodization levels necessary for beamshaping.
Pulse width variation has previously been considered for transmit apodization in association with ultrasound imaging, such as in U.S. Pat. No. 6,135,963 to Haider and U.S. Pat. No. 6,599,245 to Ma et al. As is well-known in the art, a relatively short transmit pulse is preferable for imaging, and a single pulse is often used. If a single pulse is apodized by adjusting its width, the frequency spectrum of the electrical signal supplied to the transducer to generate a corresponding ultrasound signal will likewise be significantly modified. Thus, if various transmit apodization pulse widths are applied to different array elements, the focusing, frequency spectrum and effective amplitude of the aggregate ultrasonic treatment signal emitted by the transducer will vary. This may result in undesired effects. For example, ultrasound array elements typically have limited bandwidth so there is a severe limit to the amount of useful apodization that is possible. In addition, if the transmit waveshapes are different between elements of an array, the aggregated ultrasound signals of the elements do not efficiently form a focused beam. Also, any electrical matching required to couple the electrical signal to a transducer element can drastically alter the transmit pulse shape. As a result, apodization is of limited value for ultrasound imaging.
The desired characteristics of transmit apodization of ultrasound signals for therapeutic purposes differs considerably from signals transmitted for imaging. In particular, therapeutic transducers transmit an ultrasonic pulse train for a relatively long period of time. As a result, the center of the frequency spectrum is retained generally at the pulse repetition frequency of the pulse train and is not “pulled,” rolled-off or skewed as with a transient pulse or a pulse train of short duration. If the width of the repeating pulses is short, such as a repeating pulse train of narrow spikes, the frequency spectrum broadens out and decreases in amplitude, but the peak of the spectrum remains at the pulse repetition frequency. Likewise, as the duty cycle of the pulse train increases towards a symmetrical shape (i.e., approximately equal times in “high” and “low” logical states) the effective amplitude of an array element increases. As a result of these characteristics, transmit apodization of ultrasound signals in the manner known for imaging will not produce satisfactory results when applied to therapeutic ultrasound signals.
Still, scientists and engineers continue to seek improved ultrasound medical systems. There is a need for a more efficient way to achieve transmit apodization of therapeutic ultrasound transducers.